Periodontal tissue diseases are defined as an inflammatory condition of the gum and the bone support surrounding the teeth, often leading to complications, such as to bone degradation, which in turn can impend the correct placement and fixation of dental implants after tooth loss.
Bone grafting is a surgical procedure that replaces missing bone after bone degradations and/or in order to repair bone fractures that are extremely complex, pose a significant health risk to the patient, and/or fail to heal properly. Bone grafting is currently employed to remedy bone degradation coursed by periodontal tissue diseases with varying success rates.
Bone substitutes are in general increasingly used in surgery as over two millions bone grafting procedures are performed worldwide per year. Autografts are still widely preferred for bone substitution, though the morbidity and the inherent limited availability are the main limitations. Allografts, i.e. banked bone, are osteoconductive and weakly osteoinductive, though there are still concerns about the residual infective risks, costs and donor availability issues. As an alternative, xenograft substitutes are cheaper, but their use provides contrasting results, so far. Often, bovine bone is used to replace missing bone structure; the major advantage is that larger quantities of bone can be easily acquired compared to bone of human origin. Ceramic-based synthetic bone substitutes are alternatively based on hydroxyapatite (HA) and tricalcium phosphates, and are widely used in the clinical practice. Indeed, despite being completely resorbable and weaker than cortical bone, they have exhaustively proven to be effective. Biomimetic HAs are the evolution of traditional HA and contain ions (carbonates, Si, Sr, Fl, Mg) that mimic natural HA (biomimetic HA). Injectable cements represent another evolution, enabling mininvasive techniques. Bone morphogenetic proteins (namely BMP2 and 7) are the only bone inducing growth factors approved for human use in spine surgery and for the treatment of tibial nonunion.
Most bone grafts are expected to be reabsorbed and replaced as the natural bone heals over a few months' time.
The principles involved in successful bone grafting include osteoconduction (guiding the reparative growth of the natural bone), osteoinduction (encouraging undifferentiated cells to become active osteoblasts), and osteogenesis (living bone cells in the graft material contribute to bone remodeling).
Bone graft materials are autografts, allografts, xenografts or synthetic grafts.
Autografts
Autologous (or autogenous) bone grafting involves utilizing bone obtained from the same individual receiving the graft. Bone can be harvested from non-essential bones, such as the iliac crest or the fibula, the chin, the ribs, the mandible and even parts of the skull. Autogenous bone possesses all the properties essential for bone formation: it is osteoconductive and osteoinductive, and it houses growth factors and osteogenic cells with no associated immune or infective-related risks. Autologous bone fractions are slowly replaced by newly formed host bone. The disadvantages of autografts are that a surgical donor site is required, leading to possible post-operative pain and complications. Also, they carry a likelihood of blood loss or hematomas, infection, fracture, neurovascular injury, as well as cosmetic deformity, at the explantation site, and longer operative time.
Also, autogenous bone availability in a patient represents a significant limit, especially in pediatric patients and in the elderly.
Allografts
Allograft biobanked bone represents a suitable alternative to autogenous bone, being derived from humans as well. Allograft bone can be collected from either living donors (patients total hip replacement surgery) or nonliving donors and must be processed within a bone tissue bank. Donor bone is osteoconductive, weakly osteoinductive (growth factors may still be present, depending on the processing). Also, allografts often require sterilization (gamma-irradiation), with detrimental effects on mechanical properties of bone, and deactivation of proteins normally found in healthy bone.
The limits of such transplants are costs, laborious procedure (tissue processing, harvesting), mechanical resistance (in freeze dried and irradiated), limited osteoinduction and risk of infection.
Xenografts
Xenograft bone substitutes have their origin from a species other than human, such as bovine bone (or porcine bone) which can be freeze dried or demineralized and deproteinized. Xenografts are usually only distributed as a calcified matrix. Coral based xenografts are mainly calcium carbonate (and an important proportion of fluorides, useful in the context of grafting to promote bone development) while natural human bone is made of hydroxyapatite along with calcium phosphate and carbonate. The coral material is thus either transformed industrially into hydroxyapatite through a hydrothermal process, yielding to a non-resorbable xenograft, or simply the process is omitted and the coralline material remains in its calcium carbonate state for better resorption of the graft by the natural bone. The coral xenograft is then saturated with growth enhancing gels and solutions. Xenografts have given good results in dentistry, but scarce validation is available in orthopedics.
Clinically available coral-based products are Interpore and Pro-osteon (Interpore International, Inc., Irvine, Calif.) as well as bovine derived products such as Bio-Oss (Geistlich Biomaterials, Geistlich, Switzerland), Osteograf-N (CeraMed Co., Denver, Colo.), and Endobon (Merck Co., Darmstadt, Germany).
BioOss® is a natural product with bovine origin. It is deproteinized and sintered. The material's total porosity is between 70-75% with a particle size of 250-1000 μm.
The advantages are the easy availability, the osteoconductivity, the good mechanical properties and low costs.
Synthetic Grafts
Hydroxyapatite and Tricalcium Phosphate
Generally, synthetic bone substitutes are calcium based substitutes, in particular, a mix of HA (Hydroxyapatite) and TCP (Tricalcium phosphate), HA is a relatively inert substance that is retained “in vivo” for prolonged periods of time, whereas the more porous TCP typically undergoes biodegradation within 6 weeks of its introduction into the area of bone formation. HA achieves very high mechanical strength, while TCP has lower mechanical qualities. Often the base is a biphasic calcium phosphate, which combines 40-60% TCP with 60-40% HA, yielding a more physiological balance between mechanical support and bone resorption.
Synthetic bone grafts are widely known and are proven to be safe and effective in bone substitution. HA-TCP materials are available in form of blocks, granules and injectable kits. The pore size varies between different materials but is generally within the range of 0.1 to 1000 μm, such as 100-800 μm. Pore interconnectivity is necessary for bone ingrowth. Depending on the concentration of HA and TCP, the strength is variable between 10 and 60 MP, which is lower than cortical bone compression strength (150-200 MP), which is one of the major limit of ceramic based biomaterials.
An exemplary HA-TCP material includes a porous biphasic synthetic bone-graft substitute in granulated form, herein denoted Oss. It consists of biphasic calcium phosphate, a composite of 10% hydroxyapatite and 90% β-tricalcium phosphate. The pore size is 0.1-1000 μm. The total porosity in this material is about 50-85%, such as 65±15% or 70±15%.
Another exemplary HA-TCP material is Straumann BoneCeramic® (Straumann AG, Basel, Switzerland) which is a synthetic bone-graft substitute designed for augmenting bone. It consists of biphasic calcium phosphate with a composite of 60% hydroxyapatite and 40% β-tricalcium phosphate. BoneCeramic® is 90% porous with interconnected pores of 100-500 μm in diameter.
Yet another exemplary HA-TCP material is Botiss Maxresorb® (Botiss dental GmbH, Berlin, Germany) which is a synthetic bone-graft substitute designed for augmenting bone. It consists of biphasic calcium phosphate with a composite of 60% hydroxyapatite and 40% β-tricalcium phosphate. Maxresorb® is 80% porous with interconnected pores of 200-800 μm in diameter and micropores having a diameter of 1-10 μm.
Hydroxyapatite (HA) is the primary mineral component of teeth and bone. HA ceramics come in both naturally and synthetic forms. HA and TCP ceramics are manufactured in a variety of forms including granules and porous blocks. TCP is more soluble than HA. Although HA accounts for nearly 70% of the mineral content of teeth and bone, the occurring HA in the human body exists in a substituted form. Carbonate, silicates, and magnesium among other ions, may replace hydroxyl or phosphate groups of the apatite structure. Investigators have attempted to produce HA that more closely resembles the mineral content of native bone, enhancing bioactivity and osteoconduction (Biomimetic ceramic substitutes).
Calcium Phosphate Cements
Calcium phosphate cements (CPC) are synthetic bone substitutes. The cements are a white powder, consisting of calcium phosphate, that when mixed with a liquid, forms a workable paste which can be shaped during surgery to fit the contours of bone loss. The cements harden within 20 min. The hardening reaction, which forms nanocrystalline hydroxyapatite (HA) is isothermic and occurs at physiologic pH so tissue damage does not occur during the setting reaction. CPCs were FDA approved for the treatment of non-load-bearing bone defects in 1996. HA is the primary inorganic component of natural bone which makes the hardened cement biocompatible and osteoconductive. Over time, CPCs are gradually resorbed and replaced with new bone. Because CPCs are brittle, they are used for non-load-bearing applications such as dental, crania-facial and orthopedic applications. CPCs have two significant advantages over pre-formed, sintered ceramics. First, the CPCs paste can be sculpted during surgery to fit the cavities. Second, the nanocrystalline hydroxyapatite structure of the CPC makes it osteoconductive causing it to be gradually resorbed and replaced with new bone.
Recently the research on CPC has focused on improving mechanical properties, making premixed cements, making the cement macroporous and seeding cells and growth factors into the cement.
Calcium Sulphate
Calcium sulphate (CS) is resorbed variably within 6-8 weeks. Due to rapid graft resorption, the resulting calcium-rich fluid incites inflammation. Recently many adverse or no effects were reported, mainly explained because of the too fast resorption and the production of a similar inflammatory reaction without bone formation (13-18%).
Polymer-Based Bone Graft Substitutes
Polymers have physical, mechanical, and chemical properties completely different from the other bone substitutes. The polymers can be divided into natural polymers and synthetic polymers. These, in turn, can be divided further into degradable and nondegradable types.
Degradable synthetic polymers are resorbed by the body. The benefit is that they enhance healing without remaining foreign bodies. Degradable polymers such as polylactic acid and poly(lactic-co-glycolic acid) have been used as standalone devices and as extenders of autografts and allografts.
Composite Materials
Composite of Collagen and Hydroxyapatite
Bone is mainly made of collagen (Col) and carbonate substituted hydroxyapatite (HA). Actually it is possible to obtain Col-HA by a self-assembling process on a nanometric scale.
Thus, an implant manufactured from such components is likely to behave better than other bone substitutes made as monolithic devices. Indeed, both collagen type I and hydroxyapatite were found to enhance osteoblast differentiation, but combined together, they were shown to accelerate osteogenesis.
The direct comparison of other materials compared with Col-HA composites for bone substitutes have yet to be clearly investigated. However, increasing the biomimetic properties of an implant may reduce the problems of bacterial infections associated with inserting a foreign body.
Growth Factors
Several bone-inducing growth factors are currently known in the field of the art, such as bone morphogenetic proteins (BMP), insulin growth factor (IGF), transforming growth factor (TGF), fibroblast growth factor (FGF), able to stimulate activation and migration of osteogenic stem cells and progenitor cells, and to induce revascularization.
The challenge to tissue engineers is to design and develop temporary bone scaffolds which deliver bioactive molecules and drugs or cells to the injury site and hence extend its biological functionality (accelerate healing and tissue regeneration while simultaneously preventing pathology). Although mimicking the geometric architecture of bone in a synthetic scaffold has been shown to promote favorable cellular activity, the overall capacity for a scaffold to direct cell behavior can be substantially improved through the controlled delivery of bio specific cues. Administration of growth factors and other bioactive molecules to promote bone formation and repair has achieved promising results in several preclinical and clinical models.
The efficacy of the delivery vehicle relies on its ability to provide the appropriate dose over the appropriate therapeutic time. Ideally, the presentation of bioactive molecules or drugs must be finely tuned to dynamically match the physiological needs of the tissue as it regenerates.
Many synthetic bone scaffolds rely on the delivery of single factors, which may partially explain the limited clinical utility of many current approaches. Therefore, researchers have been investigating techniques to encapsulate and release multiple bioactive molecules in a highly controlled spatial and temporal manner. Research has shown that this method significantly enhances tissue regeneration compared with the controlled release of single biological cues. The technology of incorporating multiple chemical effectors and controlling their spatial and temporal release is a very promising strategy, but is still experimental and has not yet demonstrated widespread preclinical or clinical utility to date.
The failure to identify either a single material or growth factor as the panacea for bone regeneration, or a biological scaffold that will promote integration and vascularization, has led to an increased interest in optimizing biomaterials to promote specific cell-biomaterial interactions.
New strategies work to encapsulate and release drugs which prevent pathologies that can occur post implantation of a synthetic scaffold. A wide variety of drugs have been encapsulated and released from biodegradable polymer scaffolds including antibiotics, DNA, RNA, cathepsin inhibitors, chitin, chemotherapeutics, bisphosphonates, statins, sodium fluoride, dihydropyridine, and many others. Researchers are aggressively pursuing strategies to deliver antibiotics locally to the site of injury/surgery. Although local delivery of antibiotics has a very promising outlook, there remains a number of challenges (such as antibiotic stability within the scaffold and antibiotic deactivation during fabrication), which still need to be addressed.
Emdogain®
Recent studies conducted by the present inventors have developed a gel that is injectable by means of a syringe to the site of a bone defect, named Emdogain®. This gel consists of two components, propylene glycol alginate (PGA) and Enamel Matrix Derivative (EMD). While PGA has a structural role and acts as a carrier, EMD is the active component that favours the regeneration of the diseased periodontal tissue by mediating the formation of acellular cementum at the root of the tooth and providing a foundation for the growth of the tissue associated with functional attachment. Once the gel is applied to the site of a defect, the pH tends to strive to the physiological value and when it reaches a value of 6, it causes EMD to precipitate. Afterwards, the osteoblast and cementoblast cells are enticed by the natural cocktail of isolated enamel matrix proteins to proliferate and cause the ligament extension from the gingival wall into the intrabony gum.
Several problems still remain with this approach; even if it has been shown that PGA appears to enhance the precipitation of EMD, it might also undergo a phase separation and degradation during its storage that compromises its integrity and structural role as a carrier. EMD on its own, on the other hand, has not the structural ability to sustain the charges caused by the new regenerating tissues. In general, it is required that EMD remains stable and shows a predictable evolution of properties during its sterilization and storage. Indeed, the precipitation and loss of regenerative capacity of EMD particles prior to the application has to be avoided.
Straumann® Emdogain® is a commercially available product composed of Propylene Glycol Alginate (PGA) and porcine Enamel Matrix Derivative (EMD) proteins. The PGA employed in the manufacture of Straumann® Emdogain® has a viscosity of 50-175 mPa·s (EMD in a 2% PGA aqueous solution 22° C. (Brookfield viscosity)). The composition Straumann® EMDOGAIN® itself, displays a viscosity of 3.0 Pas (3000 mPa·s at 22° C.). Upon dissolving alginates in water, the molecules hydrate and the solution gains viscosity. The viscosity of an alginate solution depends on the concentration of alginate and the length of the alginate molecules, i.e. the number of monomer units in the chains. In general, the longer the chains, the higher the viscosity at similar concentrations. The dissolved molecules are not completely flexible; rotation around the glycosidic linkages in the G-block regions is somewhat hindered, resulting in a stiffening of the chain. Solutions of stiff macromolecules are highly viscous.
It was found that because EMD is nearly 90% composed of amelogenin, it involves hierarchical structures based on nanoscale spherical agglomerates, also called nanospheres. These are generally only marginally stable and sensitive to environmental stresses such as temperature or pH changes, which may lead to an irreversible unfolding or misfolding of the proteins. This in turn leads to the proteins forming large unfolded aggregates that are unable to carry out their function as tissue healing inducer.
Although there is ample clinical evidence for periodontal regeneration following EMD application, its low viscosity is of concern. To avoid flap collapse in treatment of periodontal defects, use of EMD with a porous interconnected scaffold may be preferable. The scaffold should allow cell migration and proliferation, but also subsequently be assimilated by the surrounding. Unfortunately, amelogenin, the main active ingredient of EMD folds into a complex tertiary conformational structure and several units need to assemble into a macrostructure for the protein to assert its biological activity. This necessitates a co-ordinated aggregation and/or agglomeration on the surface of any material that is to be bioactivated, a task that has so far been hard to achieve due to the molecules easy degradation in solution or premature fall-out into disorganised sedimentary clumps before reaching the target surface.
In the present invention, a biocompatible material, developed from a natural bone mineral of bovine origin, in one embodiment available as granules of spongious bone having a bimodal pore distribution, as well as a porous synthetic bone substitute, were for the first time successfully bioactivated with a low viscosity composition comprising EMD as is described in the following.